Most sounds are transmitted in a normal ear as shown in FIG. 1 through the outer ear 101 to the tympanic membrane (eardrum) 102, which moves the bones of the middle ear 103 (malleus, incus, and stapes) that vibrate the oval window and round window openings of the cochlea 104. The cochlea 104 is a long narrow duct wound spirally about its axis for approximately two and a half turns. It includes an upper channel known as the scala vestibuli and a lower channel known as the scala tympani, which are connected by the cochlear duct. The cochlea 104 forms an upright spiraling cone with a center called the modiolar where the spiral ganglion cells of the acoustic nerve 113 reside. In response to received sounds transmitted by the middle ear 103, the fluid-filled cochlea 104 functions as a transducer to generate electric pulses which are transmitted to the cochlear nerve 113, and ultimately to the brain.
Hearing is impaired when there are problems in the ability to transduce external sounds into meaningful action potentials along the neural substrate of the cochlea 104. To improve impaired hearing, auditory prostheses have been developed. For example, when the impairment is associated with the cochlea 104, a cochlear implant with an implanted stimulation electrode can electrically stimulate auditory nerve tissue with small currents delivered by multiple electrode contacts distributed along the electrode.
FIG. 1 also shows some components of a typical cochlear implant system which includes an external microphone that provides an audio signal input to an external signal processor 111 where various signal processing schemes can be implemented. The processed signal is then converted into a digital data format, such as a sequence of data frames, for transmission into the implant 108. Besides receiving the processed audio information, the implant 108 also performs additional signal processing such as error correction, pulse formation, etc., and produces a stimulation pattern (based on the extracted audio information) that is sent through an electrode lead 109 to an implanted electrode array 110. Typically, this electrode array 110 includes multiple electrodes on its surface that provide selective stimulation of the cochlea 104.
Cochlear implant systems employ stimulation strategies that provide high-rate pulsatile stimuli in multi-channel electrode arrays. One specific example is the “Continuous Interleaved Sampling (CIS)”—strategy, as described by Wilson et al., Better Speech Recognition With Cochlear Implants, Nature, vol. 352:236-238 (1991), which is incorporated herein by reference. For CIS, symmetrical biphasic current pulses are used, which are strictly non-overlapping in time. The rate per channel typically is higher than 800 pulses/sec. Other stimulation strategies may be based on simultaneous activation of electrode currents. These approaches have proven to be successful in giving high levels of speech recognition.
Following surgical implantation, the cochlear implant (CI) must be custom fit to optimize its operation with the specific patient user. For the fitting process, it is important to know if an audible percept is elicited and how loud the percept is. Normally this information is gained using behavioral measures. For example, for each electrode contact the CI user is asked at what stimulation level the first audible percept is perceived (hearing threshold (THR)) and at what stimulation level the percept is too loud (maximum comfort level (MCL)). For CI users with limited auditory experiences or insufficient communication abilities (e.g., small children), these fitting parameters can be determined using objective measures.
One commonly used objective measure is the electrically evoked compound action potential (eCAP) which can be easily measured, but shows weak correlations with the MCL (r=0.57) and THR (r=0.55). See, for example, Miller et al., The Clinical Application Of Potentials Evoked From The Peripheral Auditory System, Hearing Research, 242(1-2), 184-197 (2008); incorporated herein by reference.
The electrically evoked stapedius reflex (eSRT) shows high correlations with the MCL. See, for example, Stephan, K. & Welzl-Müller, K., Post-Operative Stapedius Reflex Tests With Simultaneous Loudness Scaling In Patients Supplied With Cochlear Implants, Audiology, 39, 13-18 (2000) (r=0.92); and Polak, M.; Hodges, A. & Balkany, T ECAP, ESR and Subjective Levels For Two Different Nucleus 24 Electrode Arrays, Otology & Neurotology, 2005, 26, 639-645, (r=0.93-0.95); both incorporated herein by reference. But the eSRT is difficult to measure reliably, for example, movement artifacts of the impedance probe can introduce measurement artifacts.
The electrically evoked auditory brainstem response (eABR) can be measured with a slight expenditure of time, but shows only weak correlations with the MCL (r=0.54) and THR (r=0.69). See, for example, Brown et al., Relationship Between EABR Thresholds And Levels Used To Program The Clarion Speech Processor, Ann Otol Rhinol Laryngol, 108, 50-57 (1999); incorporated herein by reference.
The electrically evoked middle latency response (eMLR) also can be measured with a slight expenditure of time. But the eMLR can only be detected in 35% of awake children at initial device stimulation and 100% detectability is achieved after at least one year. See, for example, Gordon et al., Effects Of Cochlear Implant Use On The Electrically Evoked Middle Latency Response In Children, Hearing Research, 204(1-2), 78-89 (2005); incorporated herein by reference.
It is summarized in Miller et al. (2008) that “clinically useful correlations have not been observed in larger populations of CI users for either eCAP or eABR measures.” An eSRT based fitting is rarely done as it involves increased costs, for example often a second person is present to guarantee a sufficient sealed microphone probe in the CI user, especially for children. The eMLR can be measured only in a third of CI users at the initial device stimulation. U.S. Pat. No. 6,415,185 describes using a myogenic-based response for fitting CI users.